The operation of high field MRI coils (that is to say above 3 teslas) is characterised by an inhomogeneity of the radiofrequency magnetic fields transmitted or detected by a single resonator: B1+ during transmission and B1− during reception. The quantity B1+ corresponds to the circular polarization of the magnetic field turning in the same direction as the nuclear spins used for the imaging. In contrast, the quantity B1− is the polarization which turns in the opposite direction and which characterises the reception sensitivity. The inhomogeneity of magnetic field is inherent in electromagnetism equations. It increases with the frequency of the signal and consequently with the magnetic field B0.
During transmission, the inhomogeneity of B1+ results on an image in the appearance of zones of shade or artificial contrast, which are difficult to interpret. To overcome this, an array antenna formed by a multitude of resonators is used, to make uniform directly B1+ as well as the flip angle. This compensation will be all the more efficient the higher the number of resonators in an array antenna.
During reception, a larger number of resonators will provide a more uniform overall reception profile with an increase in the signal to noise ratio. This increase could be taken advantage of to enhance the resolution of the image or to reduce the acquisition time by using an acceleration method that employs the differential sensitivity between resonators due to their construction or distribution around the sample.
Conventionally, two types of resonators exist which are used in high field MRI: linear resonators (or striplines) and circular resonators or loops.
When the sample to be studied is a human head, these resonators are placed there around on a surface parallel to the head-feet axis of the patient, which also coincides with the direction of the static field B0. In this configuration, the resonators illuminate little the region of the brain situated at the top of the head. Vice-versa, they also receive little signal from this region.
Given the geometry of the head and the antennas, a first improvement solution would consist in placing, at the top of the head, a transceiver resonator having a symmetrical shape with respect to the head-feet axis, such as for example a resonator-loop of which the surface of the resonator would be orthogonal to said head-feet axis.
Nevertheless, the magnetic field radiated by a loop is essentially orthogonal to its surface. Thus, the component B1 would be essentially parallel to B0, thus inefficient in magnetic resonance. Only the near region of the conductor of the loop has a correctly oriented and efficient component B1. This observation has led to a strategy where a large number of loops of reduced size are used in arrays to seek, on the one hand, an increase in the edge effect by the increase in the total length of the conductors forming the loops and, on the other hand, a more favourable orientation of the radiated magnetic fields. Despite its relevance, this solution has two major drawbacks. Firstly, the formation and the adjustment of the array of loops are very complex. In fact, each loop has to resonate at the Larmor frequency, have a precise terminal impedance and not couple with neighbouring loops in order not to reduce the output of the array during transmission or instead increase the correlated noise during reception. Then, the array does not make it possible to explore with good sensitivity the deeper regions of the brain, especially above the thalamus because the depth of penetration of the radiated field decreases with the size of the loops.
The precession movement of the spins induces a circular polarization in the radiofrequency magnetic fields brought into play in MRI. For a given resonator, the channel that transmits the polarization B1+ will be used for exciting the spins and the channel that transmits the polarization B1− will be used to receive the relaxation signal by virtue of the reciprocity principle. Channel is taken to mean a physical port via which the resonator is connected to the outside world. Linear or circular resonators transmit a linearly polarized magnetic field. Yet, a linear polarization is the result of the superposition of two circular polarizations turning in the opposite direction, to be specific B1+ and B1−. Thus, the same channel may be used alternatively during transmission and during reception. However, half of the available power is not exploited each time. Consequently, the output and the field sensitivity of the resonator are reduced by 40% respectively during transmission and during reception.
A particularly beneficial solution consists in using a “patch” type circular resonator. Such a resonator has been used in addition to an array antenna for transmitting a circular polarization ([1] Hoffmann, J. et al. (2013), Human Brain Imaging at 9.4 T Using a Tunable Patch Antenna for Transmission. Magnetic Resonance in Medicine, 69: 1494-1500). Typically, such a resonator is constituted of a metal surface deposited on a dielectric plate of which the second face is covered by a ground plane.
This type of resonator is known and used in the field of radiofrequency antennas for data communication. The resonator described in the document WO 2007/14105 requires a disc of 320 mm diameter made of poly-tetra-fluoro-ethylene (PTFE) for a resonance frequency of 400 MHz. The diameter of the radiating element is 210 mm. The supply is realised through a 90° hybrid coupler in two points situated on the periphery of the radiating element and on two orthogonal axes going through its centre to form a transmission channel. For a “patch” resonator, the diameter of the disc is all the larger the smaller the resonance frequency. Thus for an application with an MRI scanner of 7 teslas at 298 MHz, the diameter of the radiating element according to the teaching of this document would exceed 320 mm while conserving the same substrate made of PTFE. A large diameter, besides the problem of overall dimensions in a MRI scanner, would increase the mutual coupling towards the linear or circular resonators which have to be placed as close as possible to the head and with which the “patch” resonator has to be integrated to form an array antenna.
Thus, for a medical imaging application, the external diameter of the “patch” resonator should not exceed 180 mm so that the latter can be placed at around 40 mm with respect to the top of the head.
Three methods are known for reducing the size of a “patch” resonator.
A first method consists in choosing a substrate with a markedly higher dielectric constant (typically an alumina ceramic) making it possible to reduce the wavelength of the signal propagated in the substrate. This varies in a first approximation as the inverse square root of the dielectric constant. Nevertheless, the formation of a printed circuit made of ceramic is more difficult and costly compared to a polymer substrate. It also makes the device more fragile mechanically vis-à-vis acoustic vibrations generated by the gradient magnets in a MRI scanner.
A second method consisting in making appropriately oriented slits also makes it possible to reduce the resonance frequency for a given geometry, and by extension, to reduce the size of the resonator for a given frequency ([2] Wong, K.-L. and Lin, Y.-F. (1998), Circularly polarised microstrip antenna with a tuning stub. Electronic Letters, Vol. 34, No. 9: 831-832). However, the slits such as they appear in the document have a major drawback in that they significantly reduce the magnetic field in the useful zone in MRI, close to the resonator and around the axis of symmetry.
A third method consists in mounting capacitors between the radiating element and the ground plane so as to reduce the size of a “patch” resonator. This method is particularly described in the document WO 2007/141505. Nevertheless, capacitors have intrinsic losses characterised by a quality factor. These losses increase proportionately to the value of the capacitance, and inversely proportional to the quality factor. Moreover, the capacitors should withstand voltages greater than 2000 volts for an incident power of 1 kW when the resonator is used during transmission. Yet, among available industrial capacitors, the higher the permissible voltage, the lower the quality factor. The losses reduce the output and the sensitivity of the resonator, respectively during transmission and during reception. This approach is thus only valid in the case where the output and the sensitivity are not determining criteria, as may be the case in the field of RFID (radiofrequency identification) and as is exactly the reverse case in the field of MRI.
Finally, still with the objective of reducing the overall dimensions of the resonator, it is beneficial to find a solution for obtaining a “patch” resonator without resorting to a hybrid circuit.
The reference [2] proposes a “patch” resonator without the use of a hybrid circuit. The resonator described comprises two possible connection positions, one for transmitting and receiving a LHCP polarization and the other for transmitting and receiving a RHCP polarization. The joint use of the two positions has not been envisaged for this resonator. But if the two positions had to be connected simultaneously, there would be strong mutual coupling at the level of the connection points during transmission and during reception which would result in a loss of power, around 40% of power. Such a resonator is thus not suited to be used in the field of nuclear magnetic resonance imaging.
Furthermore, the resonator described in reference [2] has irreversible adjustment means for the efficiency of the polarization and the operating frequency. These do not offer the flexibility required of a resonator used in the field of nuclear magnetic resonance imaging because these adjustments depend on the size and shape of the human head placed in its vicinity.
Given all these constraints and the prior art, “patch” resonators, although very widespread in telecommunications, are still very little used in the MRI field. In fact, none of the solutions described is satisfactory for Magnetic Resonance Imaging (MRI) at very high magnetic field.
Thus, the solutions proposed do not resolve either the problem of size of the “patches” (because they require the addition of supplementary electrical components), or even the problem of sufficient performance for an MRI application.